Medical imaging employed for cellular, pre-clinical and clinical tests and diagnosis of patients is largely classified into structural imaging and functional imaging. Structural imaging refers to structural and anatomical of the human body, and functional imaging refers to imaging of functional information of the cognitive, sensual or other functions of the human body in a direct or indirect manner. The structural or anatomical imaging technique includes computed tomography (CT), magnetic resonance imaging (MRI), or the like. For imaging of functional information about the human body's physiological and biochemical functions, positron emission tomography (PET) is widely used.
PET is a powerful biological imaging tool allowing monitoring of the functional processes of the human body in a noninvasive manner. A biological probe molecule labeled with a radioactive, positron-emitting isotope is injected into the body, and the distribution of radiation is reconstructed through tomography to visualize and quantify the physiological and biochemical responses in the body organs. The functional and molecular biological information about the brain or other organs provided by PET may be useful for the etiological study of disease, diagnosis, prognosis, monitoring after anticancer treatment, or the like.
In order to provide functional information of the human body tissues with a high sensitivity of molecular level and to overcome the problem of low resolution, PET fusion medical imaging devices such as PET-CT, PET-MRI and PET-optical imaging are being developed.
FIG. 1a is a perspective view of a PET system and FIG. 1b is an enlarged perspective view of a PET detector in the PET system.
Referring to FIGS. 1a and 1b, a PET system 100 includes a PET detector 120, a cable 130 one end of which is connected to the PET detector 120, and a PET circuitry 140 connected to the other end of the cable 130. Abed 111 for transferring a patient is provided movably on a support 113, and the PET detector 120 is provided in an imaging bore (not shown) with a tubular form configured to allow a patient to be placed therein. On one end, a plurality of scintillation crystals 121 are arranged in annular shape to form a scintillation crystal array 122. The scintillation crystal array 122 is arranged in plural numbers along a length direction of the PET detector 120. The other end of the scintillation crystals 121 is connected to a photosensor, thereby configuring the PET detector 120. The photosensor (not shown) of the PET detector 120 converts the scintillation from the scintillation crystals 121 into an electrical signal. The photosensor may be a photomultiplier tube (PMT), a positive-intrinsic-negative PIN) diode, cadmium telluride (CdTe), cadmium zinc telluride (CZT), an avalanche photodiode (APD), a Geiger-mode avalanche photodiode (GAPD), or the like.
The PMT was developed in 1974 and is currently the most frequently used photosensor in a PET device. Quantum efficiency is usually 25-30%. Recently, it was improved to 40% or higher. It is characterized by a high signal amplification factor and low noise characteristics and is advantageous in that the operation is very stable. However, high bias voltage requirement and sensitivity to magnetic field make it difficult to be applied in fusion imaging devices such as PET-MR.
Of late, semiconductor photosensors that can be used in a magnetic field were developed. Especially, the GAPD is viewed as a promising next-generation photosensor that can replace the PMT, because of high signal amplification factor, low noise characteristics, mass producibility, magnetic resonance (MR) compatibility and operability at low bias voltage. In general, the GAPD is a 4-9 mm2 sized photosensor composed of a plurality of micro-cells having a size of tens to hundreds of micrometers arranged in two dimensions. Each micro-cell consists of an APD operating in Geiger mode. When a photon is incident, an electron is emitted typically with an amplification factor of 106 times. Then, it returns to the original state as the avalanche is quenched by a quenching resistance of 100 kΩ to 1 MΩ. That is to say, each micro-cell performs on/off switch operation that detects presence or absence of photon incidence. Assuming that one gamma ray reaches the scintillation crystal and then incident on the GAPD after being converted into multiple photons, the intensity of the incident gamma ray can be calculated by counting the number of the micro-cells with on reactions.
However, for the GAPD to be utilized as the PET detector, the linearity problem for the 511 keV gamma ray has to be solved. For this reason, the GAPD consisting of many micro-cells is commonly used in spite of low quantum efficiency. Although the GAPD having a smaller number of cells per unit area and thus having high quantum efficiency can provide improved energy resolution, it is inapplicable to the PET detector because of the linearity problem for the 511 keV gamma ray.